Radiation detector and radiological image radiographing apparatus

ABSTRACT

A radiation detector and a radiological image radiographing apparatus capable of improving the quality of an obtained radiological image without causing an additional cost are provided. A first scintillator configured to include columnar crystals generating first light corresponding to a radiation emitted through a TFT substrate is laminated on the other surface of the TFT substrate that has a first photoelectric conversion element, which has one surface from which a radiation is emitted and the other surface from which at least one of the first light and the second light is emitted and which generates electric charges corresponding to the light, and a first switching element. A second scintillator which generates second light corresponding to a radiation emitted through the first scintillator and has different energy characteristics of absorbed radiations from the first scintillator is laminated on a surface of the first scintillator not facing the TFT substrate.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a Continuation of the U.S. application Ser. No.13/588,557 filed on Aug. 17, 2012, which claims priority under 35 U.S.C§119(a) to Japanese Patent Application No. 2011-185063 filed on Aug. 26,2011 and Japanese Patent Application No. 2012-167563 filed on Jul. 27,2012. Each of the above application(s) is hereby expressly incorporatedby reference, in its entirety, into the present application.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to a radiation detector and a radiologicalimage radiographing apparatus. In particular, the present inventionrelates to a radiation detector which detects an emitted radiation and aradiological image radiographing apparatus which radiographs aradiological image expressed by the radiation detected by the radiationdetector.

2. Description of the Related Art

In recent years, a radiation detector such as an FPD (Flat PanelDetector), which has a radiation-sensitive layer disposed on a TFT (ThinFilm Transistor) active matrix substrate which can convert a radiationsuch as an X-ray directly into digital data, has been put to practicaluse. A radiological image radiographing apparatus using this radiationdetector is advantageous in that an image can be immediately checked andaccordingly fluoroscopy (moving image radiographing), which is forradiographing a radiological image continuously, can be performedcompared with a radiological image radiographing apparatus using anX-ray film or an imaging plate in the related art.

As such a radiation detector, various types of radiation detectors havebeen proposed. For example, there is an indirect conversion typeradiation detector in which a radiation is first converted into light bya scintillator, such as CsI:Tl or GOS (Gd₂O₂S:Tb), and the convertedlight is converted into electric charges and stored in a sensor section,such as a photodiode. In the radiological image radiographing apparatus,the electric charges stored in the radiation detector are read as anelectric signal, and the read electric signal is amplified by anamplifier and is then converted into digital data by an A/D (analog todigital) converter.

Meanwhile, there has been a radiation detector with a phosphor layer(scintillator), which includes columnar crystals with relatively highsensitivity, in order to reduce the amount of exposure to a subject(patient).

In this technique, in order to increase the amount of radiation absorbedby the columnar crystals, it is necessary to make a scintillator layerconsiderably thick, as is also apparent from FIG. 11 in JP2008-51793A asan example. However, an increase in the thickness of the scintillatorlayer leads to an increase in cost. In addition, as the thicknessincreases, it is necessary to increase the porosity in an initialportion (base portion) of the columnar crystals. As a result, there hasbeen a problem in that the amount of emitted light in the initialportion is reduced.

That is, the diameter of a columnar portion changes with a predeterminedfluctuation during the vapor deposition of the columnar crystals.Therefore, as the thickness increases, a probability that the maximumvalue of the fluctuation will occur is increased. As a result, apossibility that columnar portions will contact each other is increased.In addition, once columnar portions contact each other, a possibilitythat the columnar portions will be fused is increased. This leads toblurring of an image. In addition, there is also a predeterminedfluctuation in the length of the columnar portion. Accordingly, if thereis adhesion of foreign matter on the substrate on which the columnarportions are vapor-deposited, the length of an abnormally grown columnarportion also increases as the thickness increases. For this reason, aprocess of reducing the length of an abnormally grown columnar portionby pressure or the like is required after the vapor deposition process.This makes the manufacturing process complicated. In addition, a normalcolumnar portion around the abnormally grown columnar portion may bedamaged due to the pressure. For this reason, when the scintillatorlayer is made thick in order to prevent the above-described fusion, itis necessary to set the filling rate of columnar crystals low (set theporosity of the initial portion high) in advance in order to prevent theabove-described fusion and to prevent the complication of the processdue to abnormal growth of columnar portions and damage to normally growncolumnar portions. For example, WO2010/007807A discloses a scintillatorin which the filling rate of columnar crystals is set to 75% to 90% whenthe thickness of the scintillator layer of the columnar crystals is 100μm to 500 μm or more. In addition, JP2006-58099A discloses ascintillator in which the filling rate of columnar crystals is set to70% to 85% when the thickness of the scintillator layer of the columnarcrystals is 500 μm or more.

As a technique which can be applied to solve the above-describedproblems, JP2002-181941A discloses a radiological digital imageradiographing apparatus that is excellent in sharpness and has highdetection efficiency. Specifically, JP2002-181941A discloses aradiological digital image radiographing apparatus which has a phosphorlayer formed of phosphor particles and binder resin and is characterizedin that the phosphor layer is configured to include a first phosphorlayer with a plate shape and a second phosphor layer which is providedin contact with the first phosphor layer and provided corresponding toeach pixel and which has an approximately columnar shape.

In addition, JP2002-181941A discloses a configuration in which theapproximately columnar second phosphor layer, the plate-shaped firstphosphor layer, and a substrate where a photoelectric conversion elementis provided are laminated sequentially from the emission side ofradiation.

Moreover, in order to provide a radiological image detector capable ofimproving the light conversion efficiency and acquiring a high-qualityimage, JP2010-121997A discloses a radiological image detector in which awavelength conversion layer including a phosphor, which receives aradiation and converts the radiation into light with a longer wavelengththan the radiation, and a detector, which detects the light converted bythe wavelength conversion layer and converts the light into an imagesignal showing a radiological image, are laminated and which ischaracterized in that the wavelength conversion layer is formed bylaminating at least two layers of a first phosphor layer and a secondphosphor layer, the second phosphor layer and the first phosphor layerare disposed in this order from the detector side, and the firstphosphor layer includes absorbent to absorb the light converted by thefirst phosphor layer.

In addition, JP2010-121997A discloses a configuration in which asubstrate where a photoelectric conversion element is provided, aplate-shaped second phosphor layer formed of GOS, and a columnar firstphosphor layer formed of CsI are laminated sequentially from theemission side of radiation.

SUMMARY OF THE INVENTION

In the technique disclosed in JP2002-181941A, however, the secondphosphor layer which has relatively high sensitivity and anapproximately columnar shape is disposed at the radiation incidenceside, but light emitted from the second phosphor layer is receivedthrough the first phosphor layer. Accordingly, there has been a problemin that the high quality is not necessarily obtained.

In addition, also in the technique disclosed in JP2010-121997A, lightfrom the first phosphor layer which has relatively high sensitivity anda columnar shape is received by the substrate through the plate-shapedsecond phosphor layer. Accordingly, similar to the technique disclosedin JP2002-181941A, there has been a problem in that the high quality isnot necessarily obtained.

The present invention has been made in view of the above-mentionedproblems and an object of the present invention is to provide aradiation detector and a radiological image radiographing apparatuscapable of improving the quality of an obtained radiological imagewithout causing an increase in cost.

In order to achieve the above-described object, according to a firstaspect of the present invention, there is provided a radiation detectorincluding: a substrate having a first photoelectric conversion element,which has one surface from which a radiation is emitted and the othersurface from which light is emitted and which generates electric chargescorresponding to the light, and a first switching element for readingthe electric charges generated by the first photoelectric conversionelement; a first phosphor layer which is laminated on the other surfaceof the substrate, generates first light corresponding to a radiationemitted through the substrate, and is configured to include columnarcrystals; and a second phosphor layer which is laminated on a surface ofthe first phosphor layer not facing the substrate, generates secondlight corresponding to a radiation emitted through the first phosphorlayer, and has different energy characteristics of absorbed radiationsfrom the first phosphor layer. Light emitted from the other surface isat least one of the first light and the second light.

In the radiation detector according to the first aspect of the presentinvention, the first phosphor layer which generates the first lightcorresponding to a radiation emitted through the substrate and isconfigured to include the columnar crystals is laminated on the othersurface of the substrate having the first photoelectric conversionelement, which has one surface from which a radiation is emitted and theother surface from which at least one of the first light and the secondlight is emitted and which generates electric charges corresponding tothe light, and the first switching element for reading the electriccharges generated by the first photoelectric conversion element. Inaddition, the second phosphor layer which generates the second lightcorresponding to a radiation emitted through the first phosphor layerand has different energy characteristics of absorbed radiations from thefirst phosphor layer is laminated on the surface of the first phosphorlayer not facing the substrate.

That is, in the present invention, since the substrate, the firstphosphor layer, and the second phosphor layer are laminated in thisorder, and a radiation is emitted from the substrate side, the surfaceof the first phosphor layer laminated on the substrate emits light morestrongly than the other surface does. Accordingly, since the lightemitting position of the first phosphor layer with respect to thesubstrate is close compared with a case where the radiation is emittedfrom the second phosphor layer side, the resolution of a radiologicalimage obtained by radiographing can be increased. As a result, thequality of the obtained radiological image can be improved.

In addition, in the present invention, the second light generated by thesecond phosphor layer is effectively guided to the substrate due to thelight guiding function by columnar crystals of the first phosphor layer.Also in this point, the quality of a radiological image can be improved.

In addition, in the present invention, a radiation which cannot beabsorbed by the first phosphor layer can be absorbed by the secondphosphor layer. Therefore, the first phosphor layer configured toinclude relatively high-cost columnar crystals can be made thin. As aresult, an increase in cost can be suppressed.

Thus, in the radiation detector according to the first aspect of thepresent invention, the first phosphor layer configured to include thecolumnar crystals which generates the first light corresponding to aradiation emitted through the substrate is laminated on the othersurface of the substrate having the first photoelectric conversionelement, which has one surface from which a radiation is emitted and theother surface from which at least one of the first light and the secondlight is emitted and which generates electric charges corresponding tothe light, and the first switching element for reading the electriccharges generated by the first photoelectric conversion element. Inaddition, the second phosphor layer which generates the second lightcorresponding to a radiation emitted through the first phosphor layerand has different energy characteristics of absorbed radiations from thefirst phosphor layer is laminated on the surface of the first phosphorlayer not facing the substrate. Therefore, the quality of an obtainedradiological image can be improved without causing an increase in cost.

Moreover, according to a second aspect of the present invention, in theradiation detector according to the first aspect of the presentinvention, the first phosphor layer may have non-columnar crystalsformed on a surface laminated on the substrate. In this case, theadhesion between the substrate and the first phosphor layer can beimproved.

Moreover, according to a third aspect of the present invention, in theradiation detector according to the first or second aspect of thepresent invention, a reflective layer may be laminated on an oppositesurface of the second phosphor layer to a surface laminated on the firstphosphor layer. In this case, light generated by each of the first andsecond phosphor layers can be efficiently condensed to the substrateside.

Moreover, according to a fourth aspect of the present invention, theradiation detector according to the first aspect of the presentinvention may further include a second substrate which is provided on anopposite surface of the second phosphor layer to a surface laminated onthe first phosphor layer and which has a second photoelectric conversionelement, which generates electric charges corresponding to the secondlight generated by the second phosphor layer, and a second switchingelement for reading the electric charges generated by the secondphotoelectric conversion element. In this case, compared with a casewhere the second substrate is not provided, light generated by thesecond phosphor layer can be used efficiently.

Moreover, according to a fifth aspect of the present invention, theradiation detector according to the first aspect of the presentinvention may further include a second substrate which is providedbetween the first and second phosphor layers and which has a secondphotoelectric conversion element, which generates electric chargescorresponding to the second light generated by the second phosphorlayer, and a second switching element for reading the electric chargesgenerated by the second photoelectric conversion element. In this case,compared with a case where the second substrate is not provided, lightgenerated by the second phosphor layer can be used efficiently.

In particular, according to a sixth aspect of the present invention, inthe radiation detector according to the fifth aspect of the presentinvention, a side of the first phosphor layer laminated on the substratemay be distal ends of the columnar crystals. In this case, compared witha case where the distal ends of the columnar crystals are laminated onthe second phosphor layer, the quality of the obtained radiologicalimage can be further improved.

Moreover, according to a seventh aspect of the present invention, in theradiation detector according to the fifth or sixth aspect of the presentinvention, a reflective layer may be laminated on an opposite surface ofthe second phosphor layer to a surface laminated on the secondsubstrate. In this case, light generated by the second phosphor layercan be efficiently condensed to the second substrate side.

Moreover, according to an eighth aspect of the present invention, in theradiation detector according to any one of the fifth to seventh aspectsof the present invention, the photoelectric conversion element of thesecond substrate may be configured to include an organic photoelectricconversion material. In this manner, noise can be effectivelysuppressed.

Moreover, according to a ninth aspect of the present invention, in theradiation detector according to any one of the fifth to eighth aspectsof the present invention, at least one of the substrate and the secondsubstrate may be a flexible substrate. In this case, even if there is arelatively large difference in the heights of the distal ends ofcolumnar crystals of the first phosphor layer, the adhesion between thesubstrate and the first phosphor layer can be improved.

Moreover, according to a tenth aspect of the present invention, in theradiation detector according to any one of the first to ninth aspects ofthe present invention, distal ends of the columnar crystals in the firstphosphor layer may be formed to be flat. In this case, the adhesionbetween the first and second phosphor layers can be improved.

Moreover, according to an eleventh aspect of the present invention, inthe radiation detector according to any one of the first to tenthaspects of the present invention, the second phosphor layer may beconfigured to include a material with a larger atomic number than anatomic number of an element which forms the columnar crystals.

Moreover, according to a twelfth aspect of the present invention, in theradiation detector according to any one of the first to eleventh aspectsof the present invention, the first phosphor layer may be configured toinclude columnar crystals of CsI, and the second phosphor layer may beconfigured to include GOS.

Moreover, according to a seventeenth aspect of the present invention,the radiation detector according to any one of the first to sixteenthaspects of the present invention may further include a buffering layerwhich is interposed between distal ends of columnar crystals and anobject to be laminated on the distal ends, is directly laminated atleast on the distal ends, and is transparent to visible light. In thiscase, even when an abnormal protrusion is generated on distal ends ofthe columnar crystals, the protrusion can be protected.

Moreover, according to a nineteenth aspect of the present invention, theradiation detector according to any one of the first to eighteenthaspects of the present invention may further include a half mirror layerwhich is interposed between the first phosphor layer and the secondphosphor layer, reflects light from the first phosphor layer, and allowsthe light to be transmitted from the second phosphor layer. In thiscase, since it is possible that the light generated in the firstphosphor layer is allowed to travel by only the columnar crystals of thefirst phosphor layer, a radiological image with less blurring can beobtained with the light guiding effect of the columnar crystals.

In addition, in order to achieve the above-described object, accordingto a twenty-first aspect of the present invention, there is provided aradiological image radiographing apparatus including: the radiationdetector according to any one of the first to twelfth aspects of thepresent invention; and a generation unit for generating imageinformation indicated by electric charges read from the radiationdetector.

In the radiological image radiographing apparatus according to thetwenty-first aspect of the present invention, image informationindicated by electric charges read from the radiation detector of thepresent invention is generated by the generation unit.

Thus, since the radiological image radiographing apparatus according tothe twenty-first aspect of the present invention includes the radiationdetector of the present invention, the quality of the obtainedradiological image can be improved without causing an increase in cost.

In addition, in order to achieve the above-described object, accordingto a twenty-third aspect of the present invention, there is provided aradiological image radiographing apparatus including: the radiationdetector according to the fourth or fifth aspect of the presentinvention; and a generation unit for generating new image information byadding, for each corresponding pixel, the image information indicated byelectric charges read from the substrate and the second substrateprovided in the radiation detector.

In the radiological image radiographing apparatus according to thetwenty-third aspect of the present invention, new image information isgenerated by adding, for each corresponding pixel, the image informationindicated by electric charges read from the substrate and the secondsubstrate provided in the radiation detector according to the fourth orfifth aspect of the present invention.

Thus, since the radiological image radiographing apparatus according tothe twenty-third aspect of the present invention includes the radiationdetector of the present invention, the quality of the obtainedradiological image can be improved without causing an increase in cost.In addition, according to the present invention, new image informationis generated by adding, for each corresponding pixel, the imageinformation indicated by electric charges read from the substrate andthe second substrate. As a result, the sensitivity of the entireradiation detector can be improved.

According to the present invention, the effect can be obtained in whichthe quality of an obtained radiological image can be improved withoutcausing an increase in cost.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a cross-sectional view showing the schematic configuration ofthree pixel units of a radiation detector according to a firstembodiment.

FIG. 2 is a schematic view showing an example of the crystalconfiguration of a scintillator according to the embodiment.

FIG. 3 is a graph showing the X-ray absorption characteristics ofvarious materials.

FIG. 4 is a cross-sectional view showing the schematic configuration ofa signal output section of one pixel unit of the radiation detectoraccording to the embodiment.

FIG. 5 is a plan view showing the configuration of the radiationdetector according to the embodiment.

FIG. 6 is a perspective view showing the configuration of an electroniccassette according to the first embodiment.

FIG. 7 is a cross-sectional view showing the configuration of theelectronic cassette according to the first embodiment.

FIG. 8 is a block diagram showing a main part configuration of theelectric system of the electronic cassette according to the firstembodiment.

FIG. 9 is a cross-sectional view showing the configuration of theradiation detector according to the first embodiment.

FIG. 10 is a cross-sectional view showing the schematic configuration ofthree pixel units of a radiation detector according to a secondembodiment.

FIG. 11 is a perspective view showing the configuration of an electroniccassette according to the second embodiment.

FIG. 12 is a cross-sectional view showing the configuration of theelectronic cassette according to the second embodiment.

FIG. 13 is a block diagram showing a main part configuration of theelectric system of the electronic cassette according to the secondembodiment.

FIG. 14 is a flow chart showing the process flow of an image informationtransmission processing program according to the second embodiment.

FIG. 15 is a cross-sectional view showing the configuration of theradiation detector according to the second embodiment.

FIG. 16 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 17 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 18 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 19 is a graph showing an example of the sensitivity characteristicsof various materials.

FIG. 20 is a graph showing an example of the sensitivity characteristicsof various materials.

FIG. 21 is a cross-sectional view providing an explanation of anabnormal protrusion generated on columnar crystals.

FIG. 22 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 23 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 24 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 25 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 26 is a cross-sectional view providing an explanation of problemswith an abnormal protrusion generated on columnar crystals.

FIG. 27 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

FIG. 28 is a cross-sectional view showing the configuration of aradiation detector according to another embodiment.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

Hereinafter, embodiments of the present invention will be described indetail with reference to the accompanying drawings.

First Embodiment

First, the configuration of an indirect conversion-type radiationdetector 20 according to the present embodiment will be described.

FIG. 1 is a schematic cross-sectional view showing the configuration ofthree pixel units of the radiation detector 20 which is an embodiment ofthe present invention.

In the radiation detector 20, a signal output section 14 (firstswitching element), a sensor section 13 (first photoelectric conversionelement), a transparent insulating layer 7, a scintillator 8A (firstphosphor layer), a scintillator 8B (second phosphor layer), a reflectivelayer 12, and a base 22 are laminated on an insulating substrate 1 inthis order. A pixel unit is formed by the signal output sections 14 andthe sensor sections 13. A plurality of pixel units are arrayed on thesubstrate 1, and each pixel unit is configured such that the signaloutput section 14 and the sensor section 13 overlap each other. Inaddition, in the present embodiment, a TFT substrate 30 is configured byforming the signal output section 14, the sensor section 13, and thetransparent insulating layer 7 in this order on the substrate 1.

The scintillator 8A is formed of columnar crystals on the sensor section13 with the transparent insulating layer 7 interposed therebetween, andis formed by depositing a phosphor which converts a radiation incidentfrom the lower side (substrate 1 side) into first light and emits thefirst light. By providing such a scintillator 8A, a radiationtransmitted through a subject is absorbed to emit light.

It is preferable that the wavelength range of the first light emittedfrom the scintillator 8A be a visible light range (wavelength of 360 nmto 830 nm). In order for the radiation detector 20 to be able to performmonochrome imaging, it is more preferable to include a green wavelengthrange.

As a phosphor used as the scintillator 8A, specifically, a phosphorincluding cesium iodide (CsI) is preferably used in the case of imagingusing an X-ray as a radiation. Especially, it is preferable to useCsI:Tl whose emission spectrum at the time of X-ray irradiation is in arage of 420 nm to 700 nm, for example. In addition, the peak emissionwavelength of CsI:Tl in the visible light range is 565 nm.

Moreover, in the present embodiment, as an example, as shown in FIG. 2,the scintillator 8A has a configuration in which a non-columnar portionformed of non-columnar crystals 71B is formed on the radiationincidence/light emission side (TFT substrate 30B side) and a columnarportion formed of columnar crystals 71A is formed on the opposite sideto the radiation incidence side of the scintillator 8A, and a materialincluding CsI is used as the scintillator 8A. By vapor-depositing thematerial directly on the TFT substrate 30, the scintillator 8A in whichthe columnar portion and the non-columnar portion are formed isobtained. In addition, in the scintillator 8A according to the presentembodiment, the average diameter of the columnar crystals 71A isapproximately uniform along the longitudinal direction of the columnarcrystals 71A.

As described above, by forming the scintillator 8A with a columnarportion, the first light generated in the scintillator 8A propagatesthrough the columnar crystals 71A and is emitted to the TFT substrate 30through the non-columnar crystals 71B. Therefore, since diffusion oflight emitted to the TFT substrate 30 side is suppressed, a decrease inthe sharpness of a radiological image obtained as a result issuppressed. In addition, the first light propagating to the distal endside of the columnar crystals 71A of the scintillator 8A is transmittedthrough the scintillator 8B and is then reflected by the reflectivelayer 12, contributing to an increase in the amount of light received bythe TFT substrate 30.

In addition, by bringing the porosity of the non-columnar portion closeto 0 (zero), reflection of light by the non-columnar portion can bepreferably suppressed. In addition, it is preferable that thenon-columnar portion be made as thin as possible (approximately 10 μm).

On the other hand, the scintillator 8B is formed so as to have differentenergy characteristics of absorbed radiations from the scintillator 8A,and is formed by depositing a phosphor which converts a radiationincident from the lower side (substrate 1 side) into second light andemits the second light. By providing such a scintillator 8B, a radiationtransmitted through the scintillator 8A is absorbed to emit light.Preferably, the wavelength range of the second light emitted from thescintillator 8B is also a visible light range.

As a phosphor used as the scintillator 8B, specifically, a phosphorincluding GOS is preferably used in the case of imaging using an X-rayas a radiation. Especially, it is preferable to use GOS:Tb. In addition,the peak emission wavelength of GOS:Tb in the visible light range is 550nm.

FIG. 3 shows the X-ray absorption characteristics of various materials.

As shown in FIG. 3, atomic numbers of elements making up the GOS arelarger than those making up the CsI. For example, in the case of GOS:Pr,a K-edge is present near 50 [KeV]. Accordingly, since the high-energyX-ray absorption rate of the GOS is higher than that of the CsI which iscolumnar crystals, a radiation which cannot be absorbed by the CsI canbe absorbed effectively. In addition, the K-edge in the GOS changes witha doping material. For example, the K-edge of GOS:Tb is approximately 60[KeV]. In addition, the atomic number referred to herein is an effectiveatomic number calculated in consideration of the composition ratio ofthe scintillator.

In addition, the reflective layer 12 reflects visible light.Accordingly, by forming the reflective layer 12, the first lightgenerated in the scintillator 8A and the second light generated in thescintillator 8B can be more efficiently guided to the sensor section 13.As a result, the sensitivity is improved. Any of a sputtering method, avapor deposition method, and a coating method may be used as a method offorming the reflective layer 12. As the reflective layer 12, it ispreferable to use materials with a high reflectance in the emissionwavelength region of the used scintillators 8A and 8B, such as Au, Ag,Cu, Al, Ni, and Ti. For example, when the scintillator 8B is GOS:Tb, itis preferable to use Ag, Al, or Cu which has a high reflectance at thewavelength of 400 to 600 nm. Regarding the thickness, the reflectance isnot obtained if the thickness is less than 0.01 μm, and the effect bythe improvement in reflectance is not obtained further even if thethickness exceeds 3 μm. Accordingly, the preferable thickness is 0.01 to3 μm.

In addition, in the present embodiment, the TFT substrate 30 is disposedon the irradiation surface side of each scintillator, and the method ofdisposing each scintillator and the TFT substrate 30 so as to satisfysuch a positional relationship is called “Irradiation Side Sampling(ISS)”. Since the radiation incidence side of the scintillator emitslight more strongly, the TFT substrate 30 and the light emittingposition of the scintillator are brought close to each other in theirradiation side sampling (ISS) in which the TFT substrate 30 isdisposed on the radiation incidence side of the scintillator, comparedwith “Penetration Side Sampling (PSS)” in which the TFT substrate 30 isdisposed on the opposite side to the radiation incidence side of thescintillator. Accordingly, the resolution of a radiological imageobtained by radiographing is high, and the amount of received light inthe TFT substrate 30 is increased. As a result, the sensitivity of theradiological image is improved.

On the other hand, the sensor section 13 has an upper electrode 6, alower electrode 2, and a photoelectric conversion layer 4 disposedbetween the upper and lower electrodes. The photoelectric conversionlayer 4 is formed of an organic photoelectric conversion material whichabsorbs the first light emitted from the scintillator 8A and the secondlight emitted from the scintillator 8B to generate electric charges.

The upper electrode 6 is preferably formed of a conductive materialwhich is transparent to at least the emission wavelength of thescintillator, since it is necessary to make the first and second lightbeams generated by the scintillators incident on the photoelectricconversion layer 4. Specifically, it is preferable to use a transparentconducting oxide (TCO) which has a high transmittance for visible lightand has a low resistance value. In addition, although a thin metal film,such as Au, may also be used as the upper electrode 6, the resistancevalue tends to increase when a transmittance of 90% or more needs to beobtained. For this reason, the TCO is preferable. For example, ITO, IZO,AZO, FTO, SnO₂, TiO₂, and ZnO₂ may be preferably used. Among these, ITOis the most preferable material from the point of view of processsimplicity, low resistance, and transparency. In addition, the upperelectrode 6 may be a common one-sheet configuration in all pixel units,or a separate upper electrode 6 may be provided in each pixel unit.

The photoelectric conversion layer 4 includes an organic photoelectricconversion material, and absorbs the first light emitted from thescintillator 8A and the second light emitted from the scintillator 8Band generates electric charges corresponding to the absorbed first andsecond light beams. The photoelectric conversion layer 4 including anorganic photoelectric conversion material as described above has anabsorption spectrum which is sharp in a visible range. Accordingly,electromagnetic waves other than the light emitted from thescintillators 8A and 8B are hardly absorbed by the photoelectricconversion layer 4. As a result, noise generated when radiations, suchas X-rays, are absorbed by the photoelectric conversion layer 4 can besuppressed effectively.

In order to absorb the first and second light beams emitted from thescintillators 8A and 8B most efficiently, it is preferable that the peakabsorption wavelength of the organic photoelectric conversion materialwhich forms the photoelectric conversion layer 4 be as close to the peakemission wavelength of each scintillator as possible. Although it isideal for the peak absorption wavelength of the organic photoelectricconversion material and the peak emission wavelength of eachscintillator to be equal, light emitted from each scintillator can besufficiently absorbed if a difference between both the wavelengths issmall. Specifically, it is preferable that the difference between thepeak absorption wavelength of the organic photoelectric conversionmaterial and the peak emission wavelength of each scintillator for theradiation be equal to or less than 10 nm. More preferably, thedifference is less than 5 nm.

As examples of the organic photoelectric conversion material which cansatisfy such conditions, a quinacridone based organic compound and aphthalocyanine based organic compound may be mentioned. For example, thepeak absorption wavelength of quinacridone in the visible light range is560 nm. Accordingly, if quinacridone is used as an organic photoelectricconversion material, CsI:Tl is used as a material of the scintillator8A, and GOS is used as a material of the scintillator 8B, the differencebetween the peak wavelengths can be made to fall within 10 nm. As aresult, the amount of charges generated in the photoelectric conversionlayer 4 can be nearly maximized.

Next, the photoelectric conversion layer 4 applicable to the radiationdetector 20 according to the present embodiment will be specificallydescribed.

An electromagnetic wave absorption/photoelectric conversion section inthe radiation detector 20 according to the present embodiment may beformed by an organic layer including a pair of electrodes 2 and 6 andthe organic photoelectric conversion layer 4 interposed between theelectrodes 2 and 6. More specifically, this organic layer may be formedby stacking or mixing of a portion which absorbs electromagnetic waves,a photoelectric conversion portion, an electron transport portion, ahole transport portion, an electron blocking portion, a hole blockingportion, a crystallization preventing portion, an electrode, aninterlayer contact improving portion, and the like.

Preferably, the above organic layer contains an organic p-type compoundor an organic n-type compound.

The organic p-type semiconductor (compound) is a semiconductor(compound) with a donor property which is mainly represented by anorganic compound with a hole transport property, and is called anorganic compound with a property prone to donating electrons. Morespecifically, the organic p-type semiconductor (compound) refers to anorganic compound with smaller ionization potential when two organicmaterials are used in a state where they are in contact with each other.Therefore, as an organic compound with a donor property, any organiccompound may be used if it is an organic compound with anelectron-donating property.

The organic n-type semiconductor (compound) is a semiconductor(compound) with an acceptor property which is mainly represented by anorganic compound with an electron transport property, and is called anorganic compound with a property of easily accepting electrons. Morespecifically, the organic n-type semiconductor (compound) refers to anorganic compound with larger electron affinity when two organicmaterials are used in a state where they are in contact with each other.Therefore, as an organic compound with an acceptor property, any organiccompound may be used if it is an organic compound with anelectron-accepting property.

Materials applicable as the organic p-type semiconductor and the organicn-type semiconductor and the configuration of the photoelectricconversion layer 4 are disclosed in detail in JP2009-32854A.Accordingly, explanation thereof will be omitted.

The thickness of the photoelectric conversion layer 4 is preferably aslarge as possible from the point of view of absorption of the firstlight from the scintillator 8A and the second light from thescintillator 8B. However, if the thickness of the photoelectricconversion layer 4 is equal to or greater than a certain value, thestrength of the electric field generated in the photoelectric conversionlayer 4 is reduced due to the bias voltage applied from both ends of thephotoelectric conversion layer 4 and as a result, it is not possible tocollect electric charges. For this reason, the thickness of thephotoelectric conversion layer 4 is preferably 30 nm or more and 300 nmor less, more preferably 50 nm or more 250 nm or less, and mostpreferably 80 nm or more and 200 nm or less.

In addition, in the radiation detector 20 shown in FIG. 1, thephotoelectric conversion layer 4 is a common one-sheet configuration inall pixel units. However, a separate photoelectric conversion layer 4may be provided in each pixel unit.

The lower electrode 2 is assumed to be a thin film divided for eachpixel unit. The lower electrode 2 may be formed of a transparent oropaque conductive material. Aluminum, silver, and the like may beappropriately used for the lower electrode 2.

The thickness of the lower electrode 2 may be set to 30 nm or more and300 nm or less, for example.

In the sensor section 13, a predetermined bias voltage may be appliedbetween the upper electrode 6 and the lower electrode 2 in order to moveone of two types of electric charges (holes and electrons) generated inthe photoelectric conversion layer 4 to the upper electrode 6 and movethe other one to the lower electrode 2. In the radiation detector 20according to the present embodiment, it is assumed that a wiring line isconnected to the upper electrode 6 and a bias voltage is applied to theupper electrode 6 through the wiring line. In addition, although thepolarity of the bias voltage is determined such that electrons generatedin the photoelectric conversion layer 4 move to the upper electrode 6and holes move to the lower electrode 2, the polarity may be reversed.

The sensor section 13 of each pixel unit may include at least the lowerelectrode 2, the photoelectric conversion layer 4, and the upperelectrode 6. However, in order to suppress an increase in a darkcurrent, it is preferable to provide at least either an electronblocking layer 3 or a hole blocking layer 5. More preferably, both theelectron blocking layer 3 and the hole blocking layer 5 are provided.

The electron blocking layer 3 can be provided between the lowerelectrode 2 and the photoelectric conversion layer 4. Accordingly, whena bias voltage is applied between the lower electrode 2 and the upperelectrode 6, a situation can be suppressed in which electrons areinjected from the lower electrode 2 to the photoelectric conversionlayer 4 and this increases a dark current.

An organic material with an electron-donating property may be used forthe electron blocking layer 3.

The material used for the electron blocking layer 3 in practice may beselected according to a material of the adjacent electrode, a materialof the adjacent photoelectric conversion layer 4, or the like.Preferably, the material used for the electron blocking layer 3 has anelectron affinity (Ea), which is larger by 1.3 eV or more than the workfunction (Wf) of the material of the adjacent electrode, and has thesame ionization potential (Ip) as the material of the adjacentphotoelectric conversion layer 4 or a smaller Ip than the material ofthe adjacent photoelectric conversion layer 4. Since materialsapplicable as the organic material with an electron-donating propertyare disclosed in detail in JP2009-32854A, explanation thereof will beomitted. In addition, the photoelectric conversion layer 4 may also beformed so as to further contain fullerene or carbon nanotubes.

In order to reliably obtain the effect of suppressing a dark current andto prevent the degradation of the photoelectric conversion efficiency ofthe sensor section 13, the thickness of the electron blocking layer 3 ispreferably 10 nm or more and 200 nm or less, more preferably 30 nm ormore and 150 nm or less, and most preferably 50 nm or more 100 nm orless.

The hole blocking layer 5 can be provided between the photoelectricconversion layer 4 and the upper electrode 6. Accordingly, when a biasvoltage is applied between the lower electrode 2 and the upper electrode6, a situation can be suppressed in which holes are injected from theupper electrode 6 to the photoelectric conversion layer 4 and thisincreases a dark current.

An organic material with an electron-accepting property may be used forthe hole blocking layer 5.

In order to reliably obtain the effect of suppressing a dark current andto prevent the degradation of the photoelectric conversion efficiency ofthe sensor section 13, the thickness of the hole blocking layer 5 ispreferably 10 nm or more and 200 nm or less, more preferably 30 nm ormore and 150 nm or less, and most preferably 50 nm or more 100 nm orless.

The material used for the hole blocking layer 5 in practice may beselected according to a material of the adjacent electrode, a materialof the adjacent photoelectric conversion layer 4, or the like.Preferably, the material used for the hole blocking layer 5 has anionization potential (Ip), which is larger by 1.3 eV or more than thework function (Wf) of the material of the adjacent electrode, and thesame electron affinity (Ea) as the material of the adjacentphotoelectric conversion layer 4 or a larger Ea than the material of theadjacent photoelectric conversion layer 4. Since materials applicable asthe organic material with an electron-accepting property are disclosedin detail in JP2009-32854A, explanation thereof will be omitted.

In addition, when a bias voltage is set such that holes of electriccharges generated in the photoelectric conversion layer 4 move to thelower electrode 2 and electrons move to the upper electrode 6, it ispreferable to reverse the positions of the electron blocking layer 3 andthe hole blocking layer 5. In addition, both the electron blocking layer3 and the hole blocking layer 5 may not be provided. If one of thelayers is provided, the effect of suppressing a dark current can beobtained to some extent.

The signal output section 14 is formed on the surface of the substrate 1below the lower electrode 2 of each pixel unit.

FIG. 4 shows the schematic configuration of the signal output section14.

A capacitor 9, which accumulates electric charges having moved to thelower electrode 2, and a field effect thin film transistor (hereinafter,simply referred to as a “thin film transistor”) 10, which converts theelectric charges accumulated in the capacitor 9 into an electric signaland outputs the electric signal, are formed corresponding to the lowerelectrode 2. The region where the capacitor 9 and the thin filmtransistor 10 are formed has a portion overlapping the lower electrode 2in plan view. By adopting such a configuration, the signal outputsection 14 and the sensor section 13 in each pixel unit overlap eachother in the thickness direction. In addition, in order to minimize theplane area of the radiation detector 20 (pixel unit), it is preferablethat the region where the capacitor 9 and the thin film transistor 10are formed be completely covered by the lower electrode 2.

The capacitor 9 is electrically connected to the corresponding lowerelectrode 2 through a wiring line of a conductive material which isformed so as to pass through an insulating layer 11 provided between thesubstrate 1 and the lower electrode 2. Accordingly, electric chargescollected in the lower electrode 2 can be moved to the capacitor 9.

The thin film transistor 10 is formed by laminating a gate electrode 15,a gate insulating layer 16, and an active layer (channel layer) 17 andforming a source electrode 18 and a drain electrode 19 further on theactive layer 17 with a predetermined distance therebetween.

For example, the active layer 17 may be formed of amorphous silicon,amorphous oxide, an organic semiconductor material, carbon nanotubes, orthe like. In addition, materials which form the active layer 17 are notlimited to these.

As amorphous oxides which can form the active layer 17, an oxidecontaining at least one of In, Ga, and Zn (for example, an In—O basedoxide) is preferably used, an oxide containing at least two of In, Ga,and Zn (for example, an In—Zn—O based oxide, an In—Ga—O based oxide, anda Ga—Zn—O based oxide) is more preferable, and an oxide containing In,Ga, and Zn is most preferable. As an In—Ga—Zn—O based amorphous oxide,an amorphous oxide whose composition in the crystalline state isexpressed as InGaO₃(ZnO)_(m) (m is a natural number of 6 or less) ispreferable. In particular, InGaZnO₄ is more preferable. In addition,materials which can form the active layer 17 are not limited to these.

As organic semiconductor materials which can form the active layer 17, aphthalocyanine compound, pentacene, vanadyl phthalocyanine, and the likemay be mentioned. However, the organic semiconductor materials which canform the active layer 17 are not limited to these. In addition, theconfiguration of the phthalocyanine compound is disclosed in detail inJP2009-212389A. Accordingly, explanation thereof will be omitted.

If the active layer 17 of the thin film transistor 10 is formed of anamorphous oxide, an organic semiconductor material, or carbon nanotubes,a radiation such as an X-ray is not absorbed or a very small amount ofradiation is absorbed even if it is absorbed. Therefore, the generationof noise in the signal output section 14 can be effectively suppressed.

In addition, when the active layer 17 is formed of carbon nanotubes, theswitching speed of the thin film transistor 10 can be increased, and thethin film transistor 10 whose light absorbance in the visible lightrange is low can be formed. In addition, when the active layer 17 isformed of carbon nanotubes, the performance of the thin film transistor10 is significantly reduced even if a very small amount of metallicimpurities are mixed into the active layer 17. Therefore, it isnecessary to form extremely high-purity carbon nanotubes by separationand extraction using centrifugation or the like.

Here, all of the amorphous oxides, the organic semiconductor materials,the carbon nanotubes, and organic photoelectric conversion materialsdescribed above may be deposited at low temperature. Therefore, as thesubstrate 1, a flexible substrate such as plastic, aramid, and abio-nano fiber may also be used without being limited to highlyheat-resistant substrates, such as a semiconductor substrate, a quartzsubstrate, and a glass substrate. Specifically, flexible substratesformed of polyester such as polyethylene terephthalate, polybutylene,and polyethylene naphthalate, polystyrene, polycarbonate, polyethersulfone, polyarylate, polyimide, polycycloolefin, norbornene resin, andpoly(chlorotrifluoroethylene), can be used. If such a flexible substrateformed of plastic is used, the weight can be reduced. This isadvantageous in carriage, for example.

In addition, an insulating layer for ensuring insulation, a gas barrierlayer for preventing the transmission of moisture or oxygen, anundercoat layer for improving the flatness or the adhesion to anelectrode, and the like may be provided on the substrate 1.

In the case of aramid, the high-temperature process at 200° or highercan be applied. Accordingly, a transparent electrode material can becured at high temperature to reduce the resistance. In addition, aramidcan allow automatic mounting of driver ICs, including the solder reflowprocess. In addition, since the thermal expansion coefficient of aramidis close to that of an ITO (indium tin oxide) or a glass substrate,there is little warping after manufacture. Accordingly, resistance tocracking is high. In addition, aramid can form a thin substrate comparedwith a glass substrate or the like. In addition, the substrate 1 mayalso be formed by laminating an ultra-thin glass substrate and aramid.

A bio-nano fiber is formed by mixing cellulose microfibril bundles(bacterial cellulose) made by bacteria (Acetobacter xylinum) with atransparent resin. The cellulose microfibril bundle has a width of 50 nmand a size equivalent to 1/10 of the visible light wavelength and alsohas high strength, high elasticity, and low thermal expansion. Byimpregnating bacterial cellulose with a transparent resin, such asacrylic resin or epoxy resin, and curing it, the bio-nano fiber whichhas an optical transmittance of approximately 90% at the wavelength of500 nm while containing 60% to 70% fibers can be obtained. Since thebio-nano fiber has a low thermal expansion coefficient (3 ppm to 7 ppm)comparable to silicon crystal, strength (460 MPa) comparable to steel,and high elasticity (30 GPa) and is also flexible, the substrate 1 whichis thinner than a glass substrate or the like can be formed.

In the meantime, in the radiation detector 20 according to the presentembodiment, the scintillator 8A is directly formed on the TFT substrate30 by vapor deposition as described above. However, the radiationdetector 20 may be manufactured by various methods without being limitedto this. Table 1 shows four examples of a method of manufacturing theradiation detector 20.

TABLE 1 Manufacturing Scintillator method Scintillator 8A Interfacestructure 8B First pattern Direct vapor Bonding Coating depositionSecond pattern Direct vapor Pressing + pouch of Coating depositionentire radiation detector Third pattern Indirect vapor Bonding orpressing Coating deposition + TFT substrate bonding, peeling off ofvapor-deposited substrate Fourth pattern Indirect vapor Coatingdeposition (vapor deposition on scintillator 8B) + TFT substrate bonding

In the manufacturing method of the first pattern, the scintillator 8A isdirectly formed on the TFT substrate 30 by vapor deposition, and thereflective layer 12 is formed on the base 22 formed of polyethyleneterephthalate or the like. Then, the scintillator 8B is formed bycoating on the reflective layer 12. Then, the surface (distal side ofcolumnar crystals) of the scintillator 8A not facing the TFT substrate30 and the surface of the scintillator 8B not facing the reflectivelayer 12 are bonded to each other using adhesive or the like.

In addition, in the manufacturing method of the second pattern, in thesame manner as in the first pattern, the scintillator 8A is directlyformed on the TFT substrate 30 by vapor deposition, and the reflectivelayer 12 is formed on the base 22 formed of polyethylene terephthalateor the like. Then, the scintillator 8B is formed by coating on thereflective layer 12. Then, pouch finishing (lamination) of the entireradiation detector 20 is performed in a state where the surface (distalside of columnar crystals) of the scintillator 8A not facing the TFTsubstrate 30 and the surface of the scintillator 8B not facing thereflective layer 12 are pressed against each other.

On the other hand, in the manufacturing method of the third pattern, thescintillator 8A is formed on a vapor-deposited substrate (not shown) byvapor deposition, and the reflective layer 12 is formed on the base 22formed of polyethylene terephthalate or the like in the same manner asin the first and second patterns. Then, the scintillator 8B is formed bycoating on the reflective layer 12. Then, the surface (distal side ofcolumnar crystals) of the scintillator 8A not facing the vapor-depositedsubstrate is bonded to the TFT substrate 30 using adhesive or the likeso that the vapor-deposited substrate is peeled off from thescintillator 8A, and the surface of the scintillator 8A not facing theTFT substrate 30 and the surface of the scintillator 8B not facing thereflective layer 12 are bonded to each other using adhesive or the likeor are pressed against each other.

In addition, in the manufacturing method of the fourth pattern, in thesame manner as in the first to third patterns, the reflective layer 12is formed on the base 22 formed of polyethylene terephthalate or thelike, and then the scintillator 8B is formed by coating on thereflective layer 12. Then, the scintillator 8A is formed on thescintillator 8B by vapor deposition, and the surface (distal side ofcolumnar crystals) of the scintillator 8A not facing the scintillator 8Bis bonded to the TFT substrate 30 using adhesive or the like. In thefourth pattern, a non-columnar portion is formed not on the TFTsubstrate 30 side but on the scintillator 8B side.

In addition, it is preferable to perform control such that the distalend of each columnar portion of the scintillator 8A is as flat aspossible. Specifically, this can be realized by controlling thetemperature of the vapor-deposited substrate at the end of vapordeposition. For example, when the temperature of the vapor-depositedsubstrate at the end of vapor deposition is set to 110°, the angle ofthe distal end is approximately 170°. When the temperature of thevapor-deposited substrate at the end of vapor deposition is set to 140°,the angle of the distal end is approximately 60°. When the temperatureof the vapor-deposited substrate at the end of vapor deposition is setto 200°, the angle of the distal end is approximately 70°. When thetemperature of the vapor-deposited substrate at the end of vapordeposition is set to 260°, the angle of the distal end is approximately120°. In addition, this control is disclosed in detail in JP2010-25620A.Accordingly, explanation thereof will be omitted.

On the other hand, as shown in FIG. 5, a plurality of pixels 32 each ofwhich is configured to include the sensor section 13, the capacitor 9,and the thin film transistor 10 are provided on the TFT substrate 30A ina two-dimensional manner in a fixed direction (row direction in FIG. 5)and a direction (column direction in FIG. 5) crossing the fixeddirection.

In addition, a plurality of gate wiring lines 34 which extend in theabove-described fixed direction (row direction) and serve to turn eachthin film transistor 10 on and off and a plurality of data wiring lines36 which extend in the above-described crossing direction (columndirection) and serve to read electric charges through the thin filmtransistor 10 in the ON state are provided in the radiation detector 20.

The radiation detector 20 has a plate shape, and has a quadrilateralshape with four sides on the outer edge in plan view. Specifically, theradiation detector 20 is formed in the rectangular shape.

Next, the configuration of a portable radiological image radiographingapparatus (hereinafter, referred to as an “electronic cassette”) 40,which radiographs a radiological image and in which the radiationdetector 20 is provided, will be described. FIG. 6 is a perspective viewshowing the configuration of the electronic cassette 40 according to thepresent embodiment.

As shown in FIG. 6, the electronic cassette 40 includes a plate-shapedhousing 41 formed of a material which allows a radiation to betransmitted therethrough. Therefore, the electronic cassette 40 has awaterproof and sealing structure. In the housing 41, the radiationdetector 20 that detects a radiation X emitted from the irradiationsurface side of the housing 41, at which the radiation X is irradiated,and transmitted through a subject and a lead (Pb) plate 43 which absorbsback scattered rays of the radiation X are disposed in this order. Inthe housing 41, a region corresponding to the arrangement position ofthe radiation detector 20 on one plate-shaped surface is a quadrilateralradiographing region 41A where a radiation can be detected. As shown inFIG. 7, the radiation detector 20 is disposed such that the TFTsubstrate 30 is located on the radiographing region 41A side, and isbonded to the inside of the housing 41 which forms the radiographingregion 41A.

In addition, at one end side of the inside of the housing 41, a case 42in which a cassette control unit 58 or a power supply unit 70 isdisposed at the position (outside the range of the radiographing region41A) not overlapping the radiation detector 20.

FIG. 8 is a block diagram showing a main part configuration of theelectric system of the electronic cassette 40 according to the presentembodiment.

In the radiation detector 20, a gate line driver 52 is disposed at oneof two adjacent sides, and a signal processing unit 54 is disposed atthe other side. Each gate wiring line 34 of the TFT substrate 30 isconnected to the gate line driver 52, and each data wiring line 36 ofthe TFT substrate 30 is connected to the signal processing unit 54.

In addition, an image memory 56, the cassette control unit 58, and aradio communication unit 60 are provided inside the housing 41.

Thin film transistors 10 of the TFT substrate 30 are sequentially turnedon in units of rows by a signal supplied through the gate wiring line 34from the gate line driver 52. Electric charges read by the thin filmtransistor 10 which has been turned on are transmitted through the datawiring line 36 as electric signals and are input to the signalprocessing unit 54. Thus, electric charges are sequentially read inunits of rows. As a result, a two-dimensional radiological image can beacquired.

Although not shown, the signal processing unit 54 has an amplifiercircuit, which amplifies an input electric signal, and a sample and holdcircuit for each data wiring line 36, and the electric signaltransmitted through each data wiring line 36 is amplified by theamplifier and is then held in the sample and hold circuit. In addition,a multiplexer and an A/D (analog to digital) converter are connected tothe output side of the sample and hold circuit in order. The electricsignal held in each sample and hold circuit is input to the multiplexerin order (serially) and is converted into digital image data by the A/Dconverter. The generation unit is included in the signal processing unit54.

The image memory 56 is connected to the signal processing unit 54, andthe image data output from the A/D converter of the signal processingunit 54 is stored in the image memory 56 in order. The image memory 56has a storage capacity capable of storing image data of a predeterminednumber of sheets. Accordingly, whenever a radiological image isradiographed, image data obtained by the radiographing is sequentiallystored in the image memory 56.

The image memory 56 is connected to the cassette control unit 58. Thecassette control unit 58 is formed by a microcomputer, and includes aCPU (Central Processing Unit) 58A, a memory 58B including a ROM (ReadOnly Memory) and a RAM (Random Access Memory), and a nonvolatile storageunit 58C such as a flash memory. The cassette control unit 58 controlsthe entire operation of the electronic cassette 40.

In addition, the radio communication unit 60 is connected to thecassette control unit 58. The radio communication unit 60 corresponds tothe wireless LAN (Local Area Network) standard represented by IEEE(Institute of Electrical and Electronics Engineers) 802.11a/b/g/n or thelike, and controls the transmission of various kinds of information toand from an external apparatus through radio communication. Through theradio communication unit 60, the cassette control unit 58 can performradio communication with an external device such as a console whichcontrols entire radiographing. Accordingly, transmission and receptionof various kinds of information between the cassette control unit 58 andthe console is possible.

In addition, the power supply unit 70 is provided in the electroniccassette 40, and various circuits or devices described above(microcomputer which functions as the gate line driver 52, the signalprocessing unit 54, the image memory 56, the radio communication unit60, or the cassette control unit 58) are operated by electric powersupplied from the power supply unit 70. The power supply unit 70 has abuilt-in battery (secondary battery which can be recharged) so as not toimpair the portability of the electronic cassette 40, and electric poweris supplied from the charged battery to various circuits or devices. Inaddition, wiring lines connecting the power supply unit 70 to variouscircuits or devices are not shown in FIG. 8.

Next, the operation of the electronic cassette 40 according to thepresent embodiment will be described.

When radiographing a radiological image, the electronic cassette 40according to the present embodiment is disposed with the radiographingregion 41A upward so as to be spaced apart from a radiation generator 80as shown in FIG. 7, and a radiographed portion B of a patient is placedon the radiographing region. The radiation generator 80 emits theradiation X of a radiation dose according to the radiographingconditions and the like given in advance. The radiation X emitted fromthe radiation generator 80 is transmitted through the radiographedportion B to carry the image information and is then irradiated to theelectronic cassette 40.

The radiation X emitted from the radiation generator 80 reaches theelectronic cassette 40 after being transmitted through the radiographedportion B. Electric charges corresponding to the dose of emittedradiation X are generated in each sensor section 13 of the radiationdetector 20 built in the electronic cassette 40, and the electriccharges generated in the sensor section 13 are accumulated in thecapacitor 9.

After the end of emission of the radiation X, the cassette control unit58 controls the gate line driver to output the ON signal from the gateline driver 52 to each gate wiring line 34 of the radiation detector 20one line at a time in order, thereby reading the image information. Theimage information read from the radiation detector 20 is stored in theimage memory 56.

Meanwhile, in the electronic cassette 40 according to the presentembodiment, as shown in FIG. 7, the radiation detector 20 is providedsuch that the radiation X is emitted from the TFT substrate 30 side.

In the radiation detector 20, as shown in FIG. 9, the scintillator 8Aconfigured to include columnar crystals is laminated on the oppositesurface of the TFT substrate 30 to the incidence side of the radiationX, and the scintillator 8B is laminated on the opposite surface of thescintillator 8A to the TFT substrate 30 side (incidence side of theradiation X).

For this reason, in the radiation detector 20, the surface of thescintillator 8A laminated on the TFT substrate 30 emits light morestrongly than the other surface does. Accordingly, since the lightemitting position of the scintillator 8A with respect to the TFTsubstrate 30 is close compared with a case where the radiation X isemitted from the scintillator 8B side, the resolution of a radiologicalimage obtained by radiographing can be increased. As a result, thequality of the obtained radiological image can be improved.

In addition, in the radiation detector 20, the second light generated bythe scintillator 8B is effectively guided to the TFT substrate 30 due tothe light guiding function by columnar crystals of the scintillator 8A.Also in this point, the quality of a radiological image can be improved.

In addition, in the radiation detector 20, a radiation which cannot beabsorbed by the scintillator 8A can be absorbed by the scintillator 8B.Therefore, the scintillator 8A configured to include relativelyhigh-cost columnar crystals can be made thin. As a result, an increasein cost can be suppressed.

In addition, in the radiation detector 20, since a non-columnar portionis provided in the scintillator 8A, the adhesion between thescintillator 8A and the TFT substrate 30 can be improved. Here, sincethe non-columnar portion is not essential, a non-columnar portion maynot be provided.

In addition, in the radiation detector 20, since the photoelectricconversion layer 4 is formed of an organic photoelectric conversionmaterial, most radiation is not absorbed in the photoelectric conversionlayer 4. For this reason, in the radiation detector 20 according to thepresent embodiment, the radiation X is transmitted through the TFTsubstrate 30 due to the ISS configuration, but the amount of radiationabsorbed by the photoelectric conversion layer 4 is small. Therefore,the deterioration of the sensitivity to the radiation X can besuppressed. In the ISS, the radiation X is transmitted through the TFTsubstrate 30 and reaches the scintillators 8A and 8B. However, when thephotoelectric conversion layer 4 of the TFT substrate 30 is formed of anorganic photoelectric conversion material, there is almost no absorptionof radiation in the photoelectric conversion layer 4 and accordingly, atleast the attenuation of the radiation X can be suppressed. This issuitable for the ISS.

In addition, both the amorphous oxide which forms the active layer 17 ofthe thin film transistor 10 and the organic photoelectric conversionmaterial which forms the photoelectric conversion layer 4 may be formedas layers at low temperature. For this reason, the substrate 1 can beformed of plastic resin, aramid, or bio-nano fiber with less absorptionof radiation. Since the substrate 1 formed in this manner absorbs asmall amount of radiation, the deterioration of the sensitivity to theradiation X can be suppressed even if a radiation is transmitted throughthe TFT substrate 30 by the ISS.

In addition, according to the present embodiment, as shown in FIG. 7,the radiation detector 20 is bonded to a portion equivalent to thephotographing region 41A in the housing 41 so that the TFT substrate 30is located on the photographing region 41A side. However, when thesubstrate 1 is formed of highly rigid plastic resin, aramid, or bio-nanofiber, the portion equivalent to the photographing region 41A of thehousing 4 can be formed to be thin since the rigidity of the radiationdetector 20 itself is high. In addition, since the radiation detector 20itself is flexible when the substrate 1 is formed of highly rigidplastic resin, aramid, or bio-nano fiber, the radiation detector 20 isdifficult to damage even if the impact is applied to the photographingregion 41A.

In addition, although the case where the transparent insulating layer 7is provided on the surface of the TFT substrate 30 on which thescintillator 8A is formed has been described in the present embodiment,the present invention is not limited to this, and the scintillator 8Amay be directly formed on the top surface of the TFT substrate 30without providing the transparent insulating layer 7.

Second Embodiment

Next, a second embodiment will be described.

First, the configuration of an indirect conversion type radiationdetector 20B according to the second embodiment will be described withreference to FIG. 10.

In the radiation detector 20B, a TFT substrate 30A obtained by forming asignal output section 14, a sensor section 13, and a transparentinsulating layer 7 in this order, a scintillator 8A, a scintillator 8B,and a TFT substrate 30B with the same configuration as the TFT substrate30A are laminated on an insulating substrate 1 in this order. A pixelunit is formed by the signal output sections 14 and the sensor sections13 of the TFT substrates 30A and 30B. A plurality of pixel units arearrayed on the substrate 1, and each pixel unit is configured such thatthe signal output section 14 and the sensor section 13 overlap eachother. In addition, the TFT substrate 30B is formed by forming thesignal output section 14 (second switching element), the sensor section13 (second photoelectric conversion element), and the transparentinsulating layer 7 on the insulating substrate 1 in this order.

In addition, since the scintillators 8A and 8B are the same as thoseprovided in the radiation detector 20 according to the first embodiment,explanation thereof will be omitted herein. In addition, since theconfigurations of the sensor section 13 and the signal output section 14are also the same as those of the sensor section 13 and the signaloutput section 14 of the radiation detector 20 according to the firstembodiment, explanation thereof will be omitted herein.

In the meantime, also in the radiation detector 20B according to thepresent embodiment, the scintillator 8A is directly formed on the TFTsubstrate 30A by vapor deposition. However, the radiation detector 20Bmay be manufactured by various methods without being limited to this.Table 2 shows four examples of a method of manufacturing the radiationdetector 20B.

TABLE 2 Manufacturing Scintillator method Scintillator 8A Interfacestructure 8B First pattern Direct vapor Bonding Coating + TFT depositionsubstrate bonding Second pattern Direct vapor Pressing + pouch Coating +TFT deposition of entire radiation substrate bonding detector Thirdpattern Indirect vapor Bonding or Coating + TFT deposition + TFTpressing substrate bonding substrate bonding, peeling off ofvapor-deposited substrate Fourth pattern Indirect vapor (Unification)Coating + TFT deposition (vapor substrate bonding deposition onscintillator 8B) + TFT substrate bonding

In the manufacturing method of the first pattern, the scintillator 8A isdirectly formed on the TFT substrate 30A by vapor deposition, and thescintillator 8B is formed by coating on the base 22 formed ofpolyethylene terephthalate or the like. Then, the surface of thescintillator 8B not facing the base 22 and the TFT substrate 30B arebonded to each other using adhesive or the like. Then, the surface(distal side of columnar crystals) of the scintillator 8A not facing theTFT substrate 30A and the surface of the scintillator 8B not facing theTFT substrate 30B are bonded to each other using adhesive or the like.

In addition, in the manufacturing method of the second pattern, in thesame manner as in the first pattern, the scintillator 8A is directlyformed on the TFT substrate 30A by vapor deposition, and thescintillator 8B is formed by coating on the base 22 formed ofpolyethylene terephthalate or the like. Then, the surface of thescintillator 8B not facing the base 22 and the TFT substrate 30B arebonded to each other using adhesive or the like. Then, pouch finishing(lamination) of the entire radiation detector 20B is performed in astate where the surface (distal side of columnar crystals) of thescintillator 8A not facing the TFT substrate 30A and the surface of thescintillator 8B not facing the TFT substrate 30B are pressed againsteach other.

On the other hand, in the manufacturing method of the third pattern, thescintillator 8A is formed on a vapor-deposited substrate (not shown) byvapor deposition, and the scintillator 8B is formed by coating on thebase 22 formed of polyethylene terephthalate or the like in the samemanner as in the first and second patterns. Then, the surface of thescintillator 8B not facing the base 22 and the TFT substrate 30B arebonded to each other using adhesive or the like. Then, the surface(distal side of columnar crystals) of the scintillator 8A not facing thevapor-deposited substrate is bonded to the TFT substrate 30A usingadhesive or the like so that the vapor-deposited substrate is peeled offfrom the scintillator 8A, and the surface of the scintillator 8A notfacing the TFT substrate 30A and the surface of the scintillator 8B notfacing the TFT substrate 30B are bonded to each other using adhesive orthe like or are pressed against each other. In the third pattern, anon-columnar portion is formed not on the TFT substrate 30A side but onthe scintillator 8B side.

In addition, in the manufacturing method of the fourth pattern, in thesame manner as in the first to third patterns, the scintillator 8B isformed by coating on the base 22 formed of polyethylene terephthalate orthe like. Then, the surface of the scintillator 8B not facing the base22 and the TFT substrate 30B are bonded to each other using adhesive orthe like. Then, the scintillator 8A is formed on the scintillator 8B byvapor deposition, and the surface (distal side of columnar crystals) ofthe scintillator 8A not facing the scintillator 8B is bonded to the TFTsubstrate 30A using adhesive or the like. Also in the fourth pattern, anon-columnar portion is formed not on the TFT substrate 30A side but onthe scintillator 8B side.

In addition, also in the radiation detector 20B according to the presentembodiment, it is preferable to perform control such that the distal endof each columnar portion of the scintillator 8A is as flat as possible.Here, “flat” means that the distal end of each columnar portion of thescintillator 8A is parallel or approximately parallel to the TFTsubstrate on which each columnar portion is formed. Specifically, thiscan be realized by controlling the temperature of the vapor-depositedsubstrate at the end of vapor deposition. For example, when thetemperature of the vapor-deposited substrate at the end of vapordeposition is set to 110°, the angle of the distal end is approximately170°. When the temperature of the vapor-deposited substrate at the endof vapor deposition is set to 140°, the angle of the distal end isapproximately 60°. When the temperature of the vapor-deposited substrateat the end of vapor deposition is set to 200°, the angle of the distalend is approximately 70°. When the temperature of the vapor-depositedsubstrate at the end of vapor deposition is set to 260°, the angle ofthe distal end is approximately 120°. In addition, this control isdisclosed in detail in JP2010-25620A. Accordingly, explanation thereofwill be omitted.

In addition, in the first to fourth patterns described above, the base22 is left on the surface of the scintillator 8B not facing the TFTsubstrate 30B. However, the base 22 may be peeled off before thescintillators 8A and 8B are bonded to each other.

In addition, similar to the TFT substrate 30 according to the firstembodiment shown in FIG. 5, a plurality of pixels 32 each of which isconfigured to include the sensor section 13, the capacitor 9, and thethin film transistor 10 are provided on the TFT substrates 30A and 30Bin a two-dimensional manner in a fixed direction (row direction) and adirection (column direction) crossing the fixed direction.

In addition, corresponding to each of the TFT substrates 30A and 30B, aplurality of gate wiring lines 34 which extend in the above-describedfixed direction (row direction) and serve to turn each thin filmtransistor 10 on and off and a plurality of data wiring lines 36 whichextend in the above-described crossing direction (column direction) andserve to read electric charges through the thin film transistor 10 inthe ON state are provided in the radiation detector 20B.

Next, the configuration of the electronic cassette 40 in which such aradiation detector 20B is provided will be described.

FIG. 11 is a perspective view showing the configuration of theelectronic cassette 40 according to the present embodiment, and FIG. 12is a cross-sectional view of the electronic cassette 40.

In the electronic cassette 40, the radiation detector 20B is disposedinside a housing 41. In the housing 41, a region corresponding to thearrangement position of the radiation detector 20B on one plate-shapedsurface is a radiographing region 41A to which a radiation is emitted atthe time of radiographing.

FIG. 13 is a block diagram showing a main part configuration of theelectric system of the electronic cassette 40 according to the presentembodiment.

In each of the TFT substrates 30A and 30B, a gate line driver 52 isdisposed at one of two adjacent sides, and a signal processing unit 54is disposed at the other side. Hereinafter, when the gate line driver 52and the signal processing unit 54 provided corresponding to the two TFTsubstrates 30A and 30B need to be distinguished from each other,reference numeral A is given to the gate line driver 52 and the signalprocessing unit 54 corresponding to the TFT substrate 30A and referencenumeral B is given to the gate line driver 52 and the signal processingunit 54 corresponding to the TFT substrate 30B in the followingexplanation.

Each gate wiring line 34 of the TFT substrate 30A is connected to thegate line driver 52A, and each data wiring line 36 of the TFT substrate30A is connected to the signal processing unit 54A. Each gate wiringline 34 of the TFT substrate 30B is connected to the gate line driver52B, and each data wiring line 36 of the TFT substrate 30B is connectedto the signal processing unit 54B.

Thin film transistors 10 of each of the TFT substrates 30A and 30B aresequentially turned on in units of rows by a signal supplied through thegate wiring line 34 from each of the gate line drivers 52A and 52B.Electric charges read by the thin film transistor 10 which has beenturned on are transmitted through the data wiring line 36 as electricsignals and are input to the signal processing units 54A and 54B. Thus,electric charges are sequentially read in units of rows. As a result, atwo-dimensional radiological image can be acquired.

The image memory 56 is connected to the signal processing units 54A and54B, and the image data output from the A/D converters of the signalprocessing units 54A and 54B is stored in the image memory 56 in order.

Since the cassette control unit 58 controls the operations of the gateline drivers 52A and 52B separately, the reading of image informationindicating a radiological image from the TFT substrates 30A and 30B canbe separately controlled.

Next, the operation of the electronic cassette 40 according to thepresent embodiment will be described.

When radiographing a radiological image, the electronic cassette 40according to the present embodiment is disposed with the radiographingregion 41A upward so as to be spaced apart from a radiation generator 80as shown in FIG. 12, and a radiographed portion B of a patient is placedin the radiographing region. The radiation generator 80 emits theradiation X of a radiation dose according to the radiographingconditions and the like given in advance. The radiation X emitted fromthe radiation generator 80 is transmitted through the radiographedportion B to carry the image information and is then irradiated to theelectronic cassette 40.

The radiation X emitted from the radiation generator 80 reaches theelectronic cassette 40 after being transmitted through the radiographedportion B. Electric charges corresponding to the dose of emittedradiation X are generated in each sensor section 13 of the radiationdetector 20B built in the electronic cassette 40, and the electriccharges generated in the sensor section 13 are accumulated in thecapacitor 9.

After the end of emission of the radiation X, the cassette control unit58 controls the gate line drivers 52A and 52B to output the ON signalfrom the gate line drivers 52A and 52B to each gate wiring line 34 ofthe TFT substrates 30A and 30B one line at a time in order, therebyreading the image information. The image information read from theradiation detector 20B is stored in the image memory 56. In addition, inthe electronic cassette 40 according to the present embodiment, imageinformation read from the TFT substrate 30A (hereinafter, referred to as“first image information”) and image information read from the TFTsubstrate 30B (hereinafter, referred to as “second image information”)are stored in different storage regions of the image memory 56.

Meanwhile, in the electronic cassette 40 according to the presentembodiment, operation mode instruction information indicating whichoperation mode is to be applied between an operation mode (hereinafter,referred to as an “additive radiographing mode”), in which the firstimage information and the second image information are transmitted froman external device such as a console that performs overall control ofthe radiation generator 80 and the electronic cassette 40 after beingadded for each corresponding pixel, and an operation mode (hereinafter,referred to as a “normal radiographing mode”), in which only the firstimage information is transmitted without performing the addition, isreceived through the radio communication unit 60. In addition, after theend of emission of the radiation X, the cassette control unit 58executes image information transmission processing for transmitting theimage information according to the operation mode indicated by theoperation mode instruction information received in advance.

Hereinafter, the operation of the electronic cassette 40 when executingthe image information transmission processing will be described withreference to FIG. 14. In addition, FIG. 14 is a flow chart showing theprocess flow of an image information transmission processing programexecuted by the CPU 58A in the cassette control unit 58 of theelectronic cassette 40 in this case, and the program is stored in thememory 58B in advance.

In step 100 in FIG. 9, it is determined whether or not the operationmode indicated by the received operation mode instruction information isthe additive radiographing mode. If positive determination is made, theprocess proceeds to step 102 and waits until both the first imageinformation and the second image information are stored in the imagememory 56.

In next step 104, the second image information is added to the firstimage information stored in the image memory 56 for each correspondingpixel. Then, the process proceeds to step 108.

On the other hand, when negative determination is made in step 100, theoperation mode indicated by the received operation mode instructioninformation is regarded as a normal radiographing mode, and the processproceeds to step 106 and waits until the first image information isstored in the image memory 56. Then, the process proceeds to step 108.

In step 108, the first image information is transmitted to the externaldevice through the radio communication unit 60. Then, this imageinformation transmission processing program ends.

Meanwhile, in the electronic cassette 40 according to the presentembodiment, as shown in FIG. 12, the radiation detector 20B is providedsuch that the radiation X is emitted from the TFT substrate 30A side.

In the radiation detector 20B, as shown in FIG. 15, the scintillator 8Aconfigured to include columnar crystals is laminated on the oppositesurface of the TFT substrate 30A to the incidence side of the radiationX, and the scintillator 8B is laminated on the opposite surface of thescintillator 8A to the TFT substrate 30A side (incidence side of theradiation X).

For this reason, in the radiation detector 20B, the surface of thescintillator 8A laminated on the TFT substrate 30A emits light morestrongly than the other surface does. Accordingly, since the lightemitting position of the scintillator 8A with respect to the TFTsubstrate 30A is close compared with a case where the radiation X isemitted from the scintillator 8B side, the resolution of a radiologicalimage obtained by radiographing can be increased. As a result, thequality of the obtained radiological image can be improved.

In addition, in the radiation detector 20B, the second light generatedby the scintillator 8B is effectively guided to the TFT substrate 30Adue to the light guiding function by columnar crystals of thescintillator 8A. Also in this point, the quality of a radiological imagecan be improved.

In addition, in the radiation detector 20B, a radiation which cannot beabsorbed by the scintillator 8A can be absorbed by the scintillator 8B.Therefore, the scintillator 8A configured to include relativelyhigh-cost columnar crystals can be made thin. As a result, an increasein cost can be suppressed.

Moreover, in the radiation detector 20B, a part of the second lightgenerated by the scintillator 8B is received in the TFT substrate 30Band accordingly image information obtained by the TFT substrate 30B canbe used. By using the result obtained by addition of this imageinformation and image information obtained by the TFT substrate 30A foreach corresponding pixel, the sensitivity of the entire radiationdetector 20B can be improved. As a result, since the dose of radiation Xemitted from the radiation generator 80 when radiographing aradiological image can be reduced, the amount of exposure to a patientcan be reduced.

For this reason, such a form of using a result obtained by adding theimage information is particularly useful in moving image radiographing.In addition, in this form, the image information obtained by thescintillator 8B is obtained by the PSS method. Therefore, theradiological image obtained by the above-described addition does notnecessarily have high quality, but sufficient quality to radiograph amoving image can be obtained. When high quality is required inradiographing a still image, it is also possible to use only the firstimage information without performing the above-described addition.

In addition, in the radiation detector 20, since a non-columnar portionis provided in the scintillator 8A, the adhesion between thescintillator 8A and the TFT substrate 30A can be improved. Here, bybringing the porosity of the non-columnar portion close to 0 (zero),reflection of light by the non-columnar portion can be preferablysuppressed. In addition, it is preferable that the non-columnar portionbe made as thin as possible (approximately 10 μm).

Thus, in the radiation detector 20B according to the present embodiment,the non-columnar portion is provided in the scintillator 8A. However,the non-columnar portion may not be provided without being limited tothis. In addition, in FIG. 15, a half mirror (not shown) may be providedbetween the base 22 and the scintillator 8A so that light from thescintillator 8A is reflected from the half mirror and is then receivedby the TFT substrate 30A and light from the scintillator 8B istransmitted through the half mirror and is then received by the TFTsubstrate 30A. Thus, since the light generated by the scintillator 8A isefficiently received by the TFT substrate 30A through the scintillator8A, the quality of the obtained radiological image can be improved.

In addition, in the radiation detector 20B, since the photoelectricconversion layer 4 is formed of an organic photoelectric conversionmaterial, most radiation is not absorbed in the photoelectric conversionlayer 4. For this reason, in the radiation detector 20B according to thepresent embodiment, the radiation X is transmitted through the TFTsubstrate 30A due to the ISS configuration, but the amount of radiationabsorbed by the photoelectric conversion layer 4 is small. Therefore,the deterioration of the sensitivity to the radiation X can besuppressed. In the ISS, the radiation X is transmitted through the TFTsubstrate 30A and reaches the scintillators 8A and 8B. However, when thephotoelectric conversion layer 4 of the TFT substrate 30A is formed ofan organic photoelectric conversion material, there is almost noabsorption of radiation in the photoelectric conversion layer 4 andaccordingly, at least the attenuation of the radiation X can besuppressed. This is suitable for the ISS.

In addition, both the amorphous oxide which forms the active layer 17 ofthe thin film transistor 10 and the organic photoelectric conversionmaterial which forms the photoelectric conversion layer 4 may be formedas layers at low temperature. For this reason, the substrate 1 can beformed of plastic resin, aramid, or bio-nano fiber with less absorptionof radiation. Since the substrate 1 formed in this manner absorbs asmall amount of radiation, the deterioration of the sensitivity to theradiation X can be suppressed even if a radiation is transmitted throughthe TFT substrate 30A by the ISS.

In addition, according to the present embodiment, as shown in FIG. 12,the radiation detector 20B is bonded to a portion equivalent to thephotographing region 41A in the housing 41 so that the TFT substrate 30Ais located on the photographing region 41A side. However, when thesubstrate 1 is formed of highly rigid plastic resin, aramid, or bio-nanofiber, the portion equivalent to the photographing region 41A of thehousing 4 can be formed to be thin since the rigidity of the radiationdetector 20B itself is high. In addition, since the radiation detector20B itself is flexible when the substrate 1 is formed of highly rigidplastic resin, aramid, or bio-nano fiber, the radiation detector 20B isdifficult to damage even if the impact is applied to the photographingregion 41A.

In addition, although the case where the transparent insulating layer 7is provided on the surface of the TFT substrate 30A on which thescintillator 8A is formed has been described in the present embodiment,the present invention is not limited to this, and the scintillator 8Amay be directly formed on the top surface of the TFT substrate 30Awithout providing the transparent insulating layer 7.

While the present invention has been described using the embodiments,the technical scope of the present invention is not limited to the scopedescribed in each embodiment described above. Various changes ormodifications may be made in the above embodiments without departingfrom the spirit and scope of the present invention, and forms in whichsuch changes or modifications are added are also included in thetechnical scope of the invention.

In addition, the above-described embodiments do not limit the inventiondefined in the appended claims, and all combinations of the featuresdescribed in the embodiments are not necessary for the solving means ofthe invention. Inventions of various stages are included in each of theembodiments described above, and various inventions may be extracted byproper combination of a plurality of components disclosed. Even if somecomponents are removed from all components shown in the embodiments, theconfiguration where some components are removed may also be extracted asan invention as long as the effect of the present invention is obtained.

For example, in each of the embodiments, the case has been described inwhich the present invention is applied to the electronic cassette 40which is a portable radiological image radiographing apparatus. However,the present invention is not limited to this, and may also be applied toa stationary radiological image radiographing apparatus.

In addition, in each of the embodiments, the case has been described inwhich a layer including CsI is applied as the first phosphor layer ofthe present invention. However, the present invention is not limited tothis, and other layers including columnar crystals, such as CsBr, can beapplied.

In addition, in each of the embodiments, the case has been described inwhich a layer including GOS is applied as the second phosphor layer ofthe present invention. However, the present invention is not limited tothis, and other phosphors such as BaFBr with different energycharacteristics of absorbed radiations from the first phosphor layer canbe applied.

In addition, although the case where the cassette control unit 58 or thepower supply unit 70 is disposed inside the housing 41 of the electroniccassette 40 so as not to overlap the case 42 and the radiation detectorhas been described in each of the embodiments, the present invention isnot limited to this. For example, the radiation detector and thecassette control unit 58 or the power supply unit 70 may be disposed soas to overlap each other.

In addition, although the case where the TFT substrate 30B is providedon the opposite surface of the scintillator 8B to the radiationincidence side has been described, the present invention is not limitedto this. For example, as shown in FIG. 16, the TFT substrate 30B mayalso be provided on the radiation incidence side surface of thescintillator 8B.

Here, in a radiation detector 20C shown in FIG. 16, a TFT substrate 30A,an adhesion layer 23, a scintillator 8A, a TFT substrate 30B, ascintillator 8B, a reflective layer 12, and a base 22 are laminated inthis order. In addition, in this radiation detector 20C, the distal endsof columnar crystals in the scintillator 8A are located on the TFTsubstrate 30A side. Therefore, the quality of a radiological imageobtained as a result can be improved.

In addition, in this form, the distal ends of the columnar crystals ofthe scintillator 8A are laminated on the TFT substrate 30A after formingthe scintillator 8A on a vapor-deposited substrate (not shown), and thenthe vapor-deposited substrate is peeled off. However, if light emittedfrom the scintillator 8A is not received by the TFT substrate 30B, theprocess of peeling off the vapor-deposited substrate is not necessary.In addition, if a light-transmissive and heat-resistant resin substrateis used as the vapor-deposited substrate, the process of peeling off thevapor-deposited substrate is not necessary either.

In addition, in the radiation detector 20C, it is preferable that boththe TFT substrates 30A and 30B be flexible substrates. In this case,even if the positions of the distal ends of columnar crystals of thescintillator 8A are not aligned, the adhesion between the scintillator8A and the TFT substrate 30A and between the scintillator 8A and the TFTsubstrate 30B can be improved. Moreover, in this case, as a flexiblesubstrate applied, it is preferable to apply a substrate, which usesultra-thin glass based on the floating method developed in recent yearsas a base, in order to improve the transmittance of radiation. Inaddition, ultra-thin glass applicable in this case is disclosed in“Success in the development of ultra-thin glass with a thickness of 0.1mm (thinnest in the world) using the floating method, Asahi Glass Co.,Ltd., [online], [Searched on Aug. 20, 2011], the Internet<URL:http://www.agc.com/news/2011/0516.pdf>”, for example.

Moreover, in the radiation detector 20C, when the photoelectricconversion layer 4 of the sensor section 13 of the TFT substrate 30B isformed of an organic photoelectric conversion material and the activelayer 17 of the thin film transistor 10 is formed of IGZO, thephotoelectric conversion layer 4 may be located on the scintillator 8Bside with respect to the thin film transistor 10 as schematically shownin FIG. 17, or the photoelectric conversion layer 4 may be located onthe scintillator 8A side with respect to the thin film transistor 10 asschematically shown in FIG. 18. In addition, when the photoelectricconversion layer 4 is located on the scintillator 8B side with respectto the thin film transistor 10, the sensitivity range of IGZO is 460 nmor less. Accordingly, since the photoelectric conversion layer 4 doesnot have sensitivity in the emission wavelength by GOS, emission by GOSdoes not become switching noise, which is preferable.

In addition, as the sensor section 13 of each of the radiation detectors20, 20B, and 20C, an organic CMOS sensor can be used in which thephotoelectric conversion layer 4 is formed of a material including anorganic photoelectric conversion material. Moreover, as the TFTsubstrates 30, 30A, and 30B of the radiation detectors 20, 20B, and 20C,an organic TFT array sheet obtained by arraying an organic transistorincluding an organic material as the thin film transistor 10 on aflexible sheet in an array form can be used. The above organic CMOSsensor is disclosed in JP2009-212377A, for example. In addition, theabove organic TFT array sheet is disclosed in “The University of Tokyohas developed the ultra-flexible organic transistor, Nihon KeizaiShimbun, [online], [Searched on May 8, 2011], the Internet<URL:http://www.nikkei.com/tech/trend/article/g=96958A9C93819499E2EAE2E0E48DE2EAE3E3E0E2E3E2E2E2E2E2E2E2;p=9694E0E7E2E6E0E2E3E2E2E0E2E0>”,for example.

When a CMOS sensor is used as the sensor section 13 of each radiationdetector, there is an advantage in that photoelectric conversion can beperformed at high speed. In addition, since the substrate can be formedto be thin, there is an advantage in that the absorption of a radiationwhen the ISS method is adopted can be suppressed and the CMOS sensor canalso be appropriately applied to photographing by mammography.

In contrast, as a defect when the CMOS sensor is used as the sensorsection 13 of each radiation detector, low resistance to radiation whena crystalline silicon substrate is used may be mentioned. For thisreason, there is also a known technique, such as providing an FOP (fiberoptic plate) between the sensor section and the TFT substrate, forexample.

In consideration of this defect, a SiC (silicon carbide) substrate maybe applied as a semiconductor substrate with high resistance toradiation. By using the SiC substrate, there is an advantage in that theISS method can be used. In addition, since SiC has low internalresistance and the small amount of heat generation compared with Si,there are advantages in that the amount of heat generation whenperforming moving image radiographing can be suppressed and asensitivity reduction according to an increase in the temperature of CsIcan be suppressed.

Thus, a substrate with high resistance to radiation, such as a SiCsubstrate, has generally a wide cap (up to approximately 3 eV). As anexample, as shown in FIG. 19, an absorption edge is approximately 440 nmcorresponding to the blue region. In this case, therefore, it is notpossible to use scintillators, such as CsI:Tl or GOS, which emits lightin the green region. In addition, FIG. 19 shows spectra of variousmaterials when quinacridone is used as an organic photoelectricconversion material.

On the other hand, since a scintillator that emits light in the greenregion has actively been studied from the sensitivity characteristics ofamorphous silicon, there is a high demand for using the scintillator.For this reason, by forming the photoelectric conversion layer 4 using amaterial including an organic photoelectric conversion material whichabsorbs light emitted in the green region, the scintillator which emitslight in the green region may be used.

When the photoelectric conversion layer 4 is formed of a materialincluding an organic photoelectric conversion material and the thin filmtransistor 10 is formed using the SiC substrate, sensitivity wavelengthregions of the photoelectric conversion layer 4 and the thin filmtransistor 10 are different. Therefore, light emitted by thescintillator does not become noise of the thin film transistor 10.

In addition, when SiC and the material including the organicphotoelectric conversion material are laminated as the photoelectricconversion layer 4, light emitted mainly in the blue region, such asCsI:Na, may be received and light emitted in the green region may alsobe received, resulting in the improvement in sensitivity. In addition,since the organic photoelectric conversion material absorbs almost noradiation, the organic photoelectric conversion material can beappropriately used for the ISS method.

In addition, the reason why SiC has high resistance to radiation is thatan atomic nucleus is not easily flipped even if struck by the radiation.This point is disclosed in “Development of a semiconductor device whichcan be used for a long time under high radiation environment such as thespace or nuclear field, the Institute of Atomic Energy Research ofJapan, [online], [Searched on May 8, 2011], the Internet <URL:http://www.jaea.go.jp/jari/jpn/publish/01/ff/ff36/sic.html>”, forexample.

In addition, as semiconductor materials with high resistance toradiation other than SiC, C (diamond), BN, GaN, AlN, ZnO, and the likemay be mentioned. The reason why these light-element semiconductormaterials have high resistance to radiation is that these are mainlywide-gap semiconductors and therefore, the reaction cross-sectional areais small since high energy is required for ionization (electron-holepair formation) and atomic displacement does not occur easily sincebonding between atoms is strong. In addition, this point is disclosed in“New development of nuclear electronics, the Institute of AdvancedElectronic Research of Japan, [online], [Searched on May 8, 2011], theInternet<URL:http://www.aist.go.jp/ETL/jp/results/bulletin/pdf/62-10to11/kobayashi150.pdf>or “Studies on radiation-proof characteristics of zinc oxide basedelectronic devices, Wakasa Wan Energy Research Center, 2009 (fiscalyear), public joint research report, March, 2010”, for example. Inaddition, the radiation-proof characteristics of GaN are disclosed in“Evaluation of radiation resistance of gallium nitride elements,University of Tohoku, [online], [Searched on May 8, 2011], the Internet<URL:http://cycgwl.cyric.tohoku.ac.jp/˜sakemi/ws2007/ws/pdf/narita.pdf”,for example.

In addition, as applications of GaN other than the blue LED, ICformation in the field of power systems has been studied since GaN has agood thermal conductivity and high insulation resistance. In addition,ZnO has been studied as an LED which emits light mainly in the blue toultraviolet region.

Meanwhile, in the case of using SiC, a band gap Eg is approximately 1.1to 2.8 eV of Si. Accordingly, the absorption wavelength λ of lightshifts to the short wavelength side. Specifically, since the wavelengthλ is 1.24/Eg×1000, the sensitivity changes at the wavelength up toapproximately 440 nm. Therefore, in the case of using SiC, as shown inFIG. 20 as an example, CsI:Na (peak wavelength: approximately 420 nm)which emits light in the blue region is appropriate as the emissionwavelength of the scintillator rather than CsI:Tl (peak wavelength:approximately 565 nm) which emits light in the green region. Since aphosphor preferably emits blue light, CsI:Na (peak wavelength:approximately 420 nm), BaFX:Eu (X is halogen such as Br or I, peakwavelength: approximately 380 nm), CaWO₄ (peak wavelength: approximately425 nm), ZnS:Ag (peak wavelength: approximately 450 nm), LaOBr:Tb,Y₂O₂S:Tb, and the like are suitable as phosphors. In particular, CsI:Na,BaFX:Eu used in the CR cassette or the like, and CaWO₄ used in a screen,a film, or the like are preferably used.

On the other hand, a CMOS sensor with high resistance to radiation maybe formed by using a structure of “Si substrate/thick film SiO₂/organicphotoelectric conversion material” based on Silicon On Insulator (SOI).

As technology applicable to this configuration, for example, “Buildingthe world's first basis for the development of high-performance logicintegrated circuit with radiation-proof characteristics by combinationof commercial state-of-the-art SOI technology and radiation-prooftechnology for space application, Space Science Laboratory, the JapanAerospace Exploration Agency (JAXA), [online], [Searched on May 8,2011], the Internet<URL:http://www.jaxa.jp/press/2010/11/20101122_soi_j.html>” may bementioned.

In addition, since the radiation resistance of the thick-film SiO₂ ishigh in the SOI, complete separation type thick-film SiO, a partialseparation type thick-film SiO, and the like may be exemplified as highradiation durable elements. In addition, these SOIs are disclosed in“Report on patent application technology trends regarding the SOI(Silicon On Insulator) technology, Japanese Patent Office, [online],[Searched on May 8, 2011], the Internet<URL:http://www.jpo.go.jp/shiryou/pdf/gidou-houkoku/soi.pdf”, forexample.

In addition, even if the thin film transistor 10 and the like of theradiation detector 20 are configured not to have light transparency (forexample, even if the active layer 17 is formed of a material with nolight transparency, such as amorphous silicon), the light-transmissiveradiation detector 20 can be obtained by disposing the thin filmtransistor 10 and the like on the light-transmissive substrate 1 (forexample, a flexible substrate formed of synthetic resin) and forming aportion of the thin film transistor 10, in which the thin filmtransistor 10 and the like are not formed, such that light istransmitted through the portion. Disposing the thin film transistor 10and the like with no light transparency on the light-transmissivesubstrate 1 can be realized by the technique of separating a fine deviceblock manufactured on a first substrate from the first substrate anddisposing the fine device block on a second substrate, specifically, byapplying an FSA (Fluidic Self-Assembly), for example. The FSA isdisclosed in “Studies on technology of self-aligned arrangement of finesemiconductor blocks, University of Toyama, [online], [Searched on May8, 2011], the Internet<URL:http://www3.u-toyama.ac.jp/maezawa/Research/FSA.html>”, forexample.

Meanwhile, as described in the above embodiments, in a case where thescintillator is configured to include the columnar crystals such as CsI,as disclosed in JP2005-148060A as an example, in some cases, aprotrusion portion 90 is formed by a partial abnormal growth of thecolumnar crystals as shown in FIG. 21 as an example, because of dust,splashing during the vapor deposition, unevenness of surface roughnessof the TFT substrate, a defect of a pinhole or a protrusion of the TFTsubstrate, and the like. Herein, numerical reference 92 in the drawingdenotes foreign matter caused by the protrusion portion 90 such as dust.As shown in FIGS. 9 and 15, when the radiation detector according to theabove embodiments is configured by forming the scintillator 8A by directvapor deposition on the TFT substrate, and when the radiation detectoris used with ISS, the boundary side of the scintillator 8B becomes theprotrusion portion 90 side.

As a solution when the protrusion portion 90 is formed as describedabove, as disclosed in JP2005-148060A, by applying a pressing force tothe distal side of the columnar crystals, the protrusion portion 90 iscrushed, the distal side is ground, or the protrusion portion 90 ismelted so that the protrusion portion 90 can be removed.

Meanwhile, as a solution other than the one described above, a method ofprotecting the protrusion 90 by providing a buffering layer on thedistal side of the scintillator 8A, without removing the protrusionportion 90 may be considered. Hereinafter, the detailed method will bedescribed.

FIG. 22 shows a configuration example of a radiation detector 20′ in acase of providing a buffering layer 94 which is transparent to visiblelight, with respect to the radiation detector 20 according to the firstembodiment. As shown in the drawing, in the radiation detector 20′, thebuffering layer 94 is interposed between the distal side of the columnarcrystals of the scintillator 8A and the scintillator 8B.

Meanwhile, FIG. 23 shows a configuration example of a radiation detector20B′ in a case of providing a buffering layer 94 with respect to theradiation detector 20B according to the second embodiment. As shown inthe drawing, in the radiation detector 20B′, the buffering layer 94 isinterposed between the distal side of the columnar crystals of thescintillator 8A and the base 22.

Further, FIG. 24 shows a configuration example of a radiation detector20 C′ in a case of providing a buffering layer 94 with respect to theradiation detector 20C shown in FIG. 16. As shown in the drawing, in theradiation detector 20C′, the buffering layer 94 is interposed betweenthe distal side of the columnar crystals of the scintillator 8A and theTFT substrate 30A.

An example of FIG. 25 is a configuration example of a radiation detector20C″ in a case of providing a buffering layer 94 when the non-columnarportion side of the scintillator 8A is configured to be the TFTsubstrate 30A side with respect to the radiation detector 20C shown inFIG. 16. As shown in the drawing, in the radiation detector 20C″, thebuffering layer 94′ is interposed between the distal side of thecolumnar crystals of the scintillator 8A and the TFT substrate 30B. Inthis case, it is not necessary for the buffering layer 94′ to betransparent to visible light.

In any case of the radiation detectors 20′, 20B′, 20C′, and 20C″, as ajoining method of the buffering layer 94 (94′) and the members on bothsurfaces thereof, a bonding method with such as an adhesive, a method ofpouch finishing (lamination) of the entire radiation detector, and thelike can be applied. As a material of the buffering layer 94 (94′), forexample, ultrathin silicon rubber (for example, thickness of 30 μm) suchas “Silius” manufactured by Fuso Rubber Co., Ltd., silicon gel, urethanegel or the like can be appropriately used.

As described above, by protecting the protrusion portion 90 by thebuffering layer, it is possible to prevent damage to the neighboringportion of the damage of the protrusion portion 90.

In the radiation detector used in a PSS in the related art, protectingthe protrusion portion was not considered other than by removing theprotrusion portion. As shown in FIG. 26 as an example, in the PPS in therelated art, it is because that, since the protrusion portion ispositioned always at the radiation incidence side, if the protrusionportion is not removed, the absorption states of the radiation of theprotrusion portion and other normal portions greatly changes, and as aresult, image defects are generated.

Meanwhile, in a case where the radiation detector according theembodiment is configured by forming the scintillator by the direct vapordeposition on the TFT substrate and in a case of using the radiationdetector in ISS, since the protrusion portion is not positioned at theradiation incidence side, the change of radiation absorption states asdescribed above is small so that it is not necessary to remove theprotrusion portion.

As described in the embodiments, when the scintillator is configured toinclude the columnar crystals such as CsI, allowing the light generatedin the scintillator to travel as much as possible in the columnarcrystals of the scintillator makes it possible to obtain a radiologicalimage with less blurring with the light guiding effect of the columnarcrystals.

As shown in FIGS. 27 and 28 as an example, it is preferable that theradiation detector described in the embodiments reflect light from thescintillator 8A and the half mirror layer 96 which transmits light fromthe scintillator 8B be interposed between the scintillator 8A and thescintillator 8B. FIG. 27 is a configuration example of a case where thehalf mirror layer 96 is interposed with respect to the radiationdetector 20 according to the first embodiment. FIG. 28 is aconfiguration example of a case where the half mirror layer 96 isinterposed with respect to the radiation detector 20B according to thesecond embodiment.

In any case, a joining method of the half mirror layer 96 and themembers on both surfaces thereof, a bonding method with such anadhesive, a method of pouch finishing (lamination) of the entireradiation detector, and the like can be applied.

As a material of the half mirror layer 96, for example, metals such asAg, Al, NiAl or the like can be applied. A metal layer having athickness of 2 nm or more and 100 nm or less can be applied. A method ofpreparing the half mirror layer 96 and the like is disclosed in“Industrial Research Institute of Niigata Prefecture, Technology andResearch Report 2005 No. 34, “Study of New Functional Thin Films”,[online], [Searched on Jul. 22, 2012], the Internet<URL:http://www.iri.pref.niigata.jp/pdf/houkoku/h16.pdf”, for example.Accordingly, explanation thereof will be omitted.

What is claimed is:
 1. A radiation detector comprising: a firstsubstrate having a first photoelectric conversion element, which has onesurface from which a radiation is emitted and the other surface fromwhich light is emitted and which generates electric chargescorresponding to the light, and a first switching element for readingthe electric charges generated by the first photoelectric conversionelement; a first phosphor layer which is laminated on the other surfaceof the first substrate, generates first light corresponding to aradiation emitted through the first substrate, is configured to includecolumnar crystals, and has non-columnar crystals formed on a surfacelaminated on the first substrate wherein distal end of the columnarcrystals is positioned at an opposite side to the first substrate side;a second phosphor layer which is laminated on a surface of the firstphosphor layer not facing the first substrate, generates second lightcorresponding to a radiation emitted through the first phosphor layer,and has different energy characteristics of absorbed radiations from thefirst phosphor layer; a second substrate which is provided between thefirst phosphor layer and the second phosphor layer and which has asecond photoelectric conversion element, which generates electriccharges corresponding to the second light generated by the secondphosphor layer, and a second switching element for reading the electriccharges generated by the second photoelectric conversion element; and abuffering layer which is interposed between the distal ends of thecolumnar crystals and the second substrate, abutting at least at thedistal ends of the columnar crystals to protect the distal end of thecolumnar crystals, wherein light emitted from the other surface is atleast one of the first light and the second light.
 2. The radiationdetector according to claim 1, wherein the buffering layer isnon-transparent to visible light.
 3. The radiation detector according toclaim 1, wherein at least a part of the distal end of the columnarcrystals has a protrusion portion protruded toward the buffering layerside, and the buffering layer protects the distal end having protrusionportion.
 4. The radiation detector according to claim 2, wherein atleast a part of the distal end of the columnar crystals has a protrusionportion protruded toward the buffering layer side, and the bufferinglayer protects the distal end having protrusion portion.
 5. Theradiation detector according to claim 1, wherein the buffering layer haselasticity.
 6. The radiation detector according to claim 2, wherein thebuffering layer has elasticity.
 7. The radiation detector according toclaim 3, wherein the buffering layer has elasticity.
 8. The radiationdetector according to claim 4, wherein the buffering layer haselasticity.
 9. The radiation detector according to claim 5, wherein thebuffering layer is made from at least one of silicon rubber, silicon geland urethane gel.
 10. The radiation detector according to claim 6,wherein the buffering layer is made from at least one of silicon rubber,silicon gel and urethane gel.
 11. The radiation detector according toclaim 3, wherein a thickness of the buffering layer in a laminatingdirection that the first substrate, the first phosphor layer, thebuffering layer, the second substrate and the second phosphor layer arelaminated is larger than a length of the protrusion portion in thelaminating direction.
 12. The radiation detector according to claim 4,wherein a thickness of the buffering layer in a laminating directionthat the first substrate, the first phosphor layer, the buffering layer,the second substrate and the second phosphor layer are laminated islarger than a length of the protrusion portion in the laminatingdirection.
 13. The radiation detector according to claim 1, wherein anouter diameter of the buffering layer is larger than an outer diameterof the first phosphor layer in a cross section along a laminatingdirection that the first substrate, the first phosphor layer, thebuffering layer, the second substrate and the second phosphor layer arelaminated.
 14. The radiation detector according to claim 2, wherein anouter diameter of the buffering layer is larger than an outer diameterof the first phosphor layer in a cross section along a laminatingdirection that the first substrate, the first phosphor layer, thebuffering layer, the second substrate and the second phosphor layer arelaminated.
 15. The radiation detector according to claim 1, wherein thesecond phosphor layer has a plate-shape.
 16. The radiation detectoraccording to claim 1, wherein the second phosphor layer is configured toinclude GOS.
 17. The radiation detector according to claim 1, wherein anabsorption rate of a low-energy radiation of the first phosphor layer ishigher than an absorption rate of a low-energy radiation of the secondphosphor layer, and an absorption rate of a high-energy radiation of thesecond phosphor layer is higher than an absorption rate of a high-energyradiation the first phosphor layer.
 18. The radiation detector accordingto claim 17, wherein a radiation energy of the low-energy radiation islower than K-edge, and a radiation energy of the high-energy radiationis higher than K-edge.
 19. A radiological image radiographing apparatuscomprising: the radiation detector according to claim 1; and ageneration unit that generates image information indicated by electriccharges read from the radiation detector.
 20. A radiological imageradiographing apparatus comprising: the radiation detector according toclaim 1; and a generation unit that generates new image information byadding, for each corresponding pixel, the image information indicated byelectric charges read from the first substrate and the second substrateprovided in the radiation detector.